Magnetic resonance imaging (MRI) is a medical imaging modality that can create images of the inside of a human body without using x-rays or other ionizing radiation. MRI uses a powerful magnet to create a strong, uniform, static magnetic field (i.e., the “main magnetic field”). When a human body, or part of a human body, is placed in the main magnetic field, the nuclear spins that are associated with the hydrogen nuclei in tissue water or fat become polarized. This means that the magnetic moments that are associated with these spins become preferentially aligned along the direction of the main magnetic field, resulting in a small net tissue magnetization along that axis (the “z axis,” by convention). An MRI system also comprises components called gradient coils that produce smaller amplitude, spatially varying magnetic fields when a current is applied to them. Typically, gradient coils are designed to produce a magnetic field component that is aligned along the z axis and that varies linearly in amplitude with position along one of the x, y or z axes. The effect of a gradient coil is to create a small ramp on the magnetic field strength and concomitantly on the resonant frequency of the nuclear spins, along a single axis. Three gradient coils with orthogonal axes are used to “spatially encode” the MR signal by creating a signature resonance frequency at each location in the body. Radio frequency (RF) coils are used to create pulses of RF energy at or near the resonance frequency of the hydrogen nuclei. The RF coils are used to add energy to the nuclear spin system in a controlled fashion. As the nuclear spins then relax back to their rest energy state, they give up energy in the form of an RF signal. This signal is detected by the MRI system and is transformed into an image using a computer and known reconstruction algorithms.
MRI systems require a uniform main magnetic field, B0, in the imaging volume, however, inhomogeneities in the magnetic field may be introduced by various factors such as manufacturing tolerances, environmental effects, design restrictions, imperfections in the magnet, ferromagnetic material near the installation site, and so forth. Inhomogeneities in the magnetic field, B0, can adversely affect data acquisition and reconstruction of an MR image. For example, magnetic field inhomogeneities may distort position information in the scan volume and degrade the image quality. A process known as “shimming” may be used to compensate for or remove inhomogeneities from the magnetic field, B0. An MRI magnet may be shimmed using shim or correction coils (active shimming) or passive shims such as pieces of ferromagnetic materials (passive shimming).
Active shimming uses dedicated coils in the magnet to generate a corrective magnetic field. Typically, a current is passed through the shim coils to create the corrective magnetic fields. The current through the shim coils may be adjusted or regulated to provide the appropriate corrective field. Shim coils may be resistive, superconducting or a combination of both. Superconducting shim coils are located inside the magnet and operate in a helium environment. Superconducting shim coils are used to compensate the inhomogeneities (harmonics) caused either by manufacturing tolerances or by the magnetic environment of the scanning room. Typically, the current in the superconducting shim coils is adjusted to a proper value(s) during installation or maintenance of the MRI scanner. Once the current is adjusted to the proper value(s), the current values are fixed and the superconducting coils operate in a persistent mode. To provide static compensation of patient-induced harmonics, which may vary from scan to scan, resistive shim coils (so-called high order shim coils) may be used. The resistive shim coils are often incorporated in the gradient assembly of an MRI scanner and typically include a second order set of shim coils for which the current may be adjusted between scans.
The development of more advanced MRI imaging techniques has created increasingly tight targets for homogeneity of the main magnetic field, B0, during imaging of a patient, for example, when using protocols such as fMRI (functional MRI). In addition, the increasing size of the main magnetic field, B0, in modern MRI scanners results in more stringent shimming requirements. Dynamic shimming may be used to address the more stringent shimming and homogeneity requirements. Dynamic shimming performs real-time compensation of field distortions (for example, the field distortions created by non-uniform distribution of magnetic susceptibility of a patient) that varies with motion such as during patient breathing. Resistive shim coils may be used to compensate for time varying harmonics by feeding the resistive shim coils with varying currents. Using the existing high order resistive shim coil set for dynamic shimming, however, may present a number of problems because the coils are unshielded. For example, if the unshielded resistive shim coils are pulsed, the resistive shim coils may create substantial eddy currents in the magnet structure that decay with varying time constants. The eddy currents may compromise image quality and may be difficult to compensate or correct. Coupling of the unshielded shim coils with the magnet structure, circuits and coils may create additional problems for dynamic shimming.
To address such problems, a shielded resistive shim coil set may be built for dynamic shimming, with its stray field minimized to reduce the potential interaction with the magnet structure and circuits. A shielded resistive shim coil set may be located in the gradient coil assembly. This, however, may significantly affect the cost, complexity and reliability of the gradient coil assembly. For example, an additional layer of shielding resistive shield coils has to be placed within a space inside the gradient coil assembly that is already occupied by other elements. The necessary radial separation between the main and shield shim coils to provide the required effectiveness (which increases with the radial distance between the main and shielding coil sets) may also be difficult to achieve within the limited radial space of the gradient coil assembly. Also, the effectiveness of the main shim coils is diminished during static compensation due to the opposite fields generated by the shielding shim coils. Since the main and shielding shim coils are connected in series, the same current is applied during all operations which can result in reduction of shim strength of the static compensation component. In addition, dynamic shimming capabilities may only be required at a limited number of sites with applications for which these advanced capabilities are critical. Incorporating a shielded high order resistive shim set in a standard gradient coil assembly would impose an unnecessary cost burden and reduction of static shimming capabilities on scanners not requiring the dynamic shimming.
There is a need for a system and method for shielded dynamic shimming that may be used with existing MRI scanners and that does not compromise the performance or cost of the static shimming components. It would be advantageous to provide a shielded dynamic shim system that utilizes a dynamic shim insert and a high order shim coil set in an MRI scanner.